Gamma ray imaging is currently used in medicine to obtain 3D images of patients' internal organs. Positron Emission Tomography (PET) is a medical gamma ray imaging technique frequently used for this purpose. FIGS. 2A, 2B and 3 show representative prior art systems. Prior to an imaging procedure, a patient is given a radiopharmaceutical, which contains a positron emitting substance and which is selectively accumulated in a region of interest. When a positron emitted by the radiopharmaceutical encounters an electron, the electron-positron pair annihilates, emitting two gamma photons of 511 keV each, flying in opposite directions. The simultaneous detection of these two 511 keV gamma photons by two gamma detectors 40 positioned opposite to each other (as shown in FIG. 1), indicates that a positron has been emitted and annihilated inside an organ of a patient 500. The simultaneous attribution of 2D coordinates to each one of the photons allows for the determination of the photon's line of flight. The position of the annihilation is along this line. When a multitude of gamma photon pairs are detected and the information processed using appropriate algorithms, electronic circuitry, software, etc., a 3D image of the organ under examination is reconstructed.
Further and more detailed descriptions and analysis of PET will be found in "Performance Parameters of a Positron Imaging Camera", by Gerd Muehllehner et al., IEEE Transactions of Nuclear Science, Volume NS-23, No. 1, February 1976 and in "Performance Parameters of a Longitudinal Tomographic Positron Imaging System," by Paans et al., Nuclear Instruments and Methods, Volume 192, Nos. 2, 3, pages 491 -500, Feb. 1, 1982, the disclosures of which are incorporated herein by reference.
Low energy, stray, gamma photons, resulting from 511 keV gamma photons scattered within patient's body, are also present during coincidence measurements and may also reach one or both detectors. These scattered low energy gamma photons do not contain any usable and/or valid information. If these stray gamma photons are allowed to reach the detectors, they increase the count rate at the detector while not adding any usable information. These additional counts, while they may be rejected later, reduce the ability of the detector to detect "real" events, at a high rate.
Another problem encountered in coincidence gamma imaging concerns attenuation artifacts caused by absorption by the patient body and scattering. In order to correct for these effects, a 3D distribution of patient's absorption is preferably previously measured.
Attenuation may be measured (see FIG. 2A) by scanning patient 500 using a collimated line source 98 situated opposite a collimated detector 40 or two collimated line sources 98, opposite two collimated perpendicular detectors 40 and 40'. When attenuation and coincidence measurements are to be performed consecutively, the configuration of the apparatus has to be changed. This procedure is very time consuming and cumbersome for the following reasons:
a) During coincidence measurements, detectors 40 are positioned parallel to each other (see FIG. 2B). PA1 b) In order to improve the resolution of the attenuation measurements, a collimator 54 is used on the detector side (see FIG. 2A). Coincidence measurements use no collimators on detectors (see FIG. 2B); PA1 c) A "Filter" 56 (see FIGS. 2B), used in coincidence measurements contains a graded absorber 58 that selectively absorbs, and thus, protects the detectors from large flux of low energy, scattered, stray gamma photons. A line source 98 used in attenuation measurements is, for practical reasons, a source emitting low energy gamma photons (e.g., 100 keV Gd 153). These gamma photons cannot penetrate graded absorber 58. PA1 if a collimation width 62 is larger than source 98 diameter 66, no substantial collimation exists (see FIG. 4A); PA1 if collimation width 62 is substantially the same as line source diameter 66, sensitivity related to manufacturing tolerances is maximal and non-uniform radiation is generated (see FIG. 4B); and PA1 if line source diameter 66 is larger than collimation width 62, and a loss of potential radioactivity 68 results. The smaller the ratio of the width of the slit to its length in the direction of the rays, the better the collimation and the greater the loss. PA1 a) The detector intrinsic resolution, i.e., the ability of the detector to accurately determine location 72 of interaction of a gamma photon 70, with scintillation crystal 42. The thicker the crystal, the higher the probability that the photon interacts with the crystal. However, as is evident from FIGS. 5A and 5B, the detector intrinsic resolution is reduced with increasing thickness. In the absence of accuracy in depth discrimination, gamma photons 70 are assumed to interact with the scintillation crystal at its median 84; PA1 b) The accuracy with which the gantry position is determined; and PA1 c) Loss of resolution due to reconstruction algorithms. PA1 a. a gamma ray collimator assembly having a first collimator portion and a second collimator portion, said first and second portions having different acceptance angles and wherein the collimator portions are formed by spaced openings and wherein the septa openings are different for the two collimator portions; and PA1 b. a gamma ray detector wherein said gamma ray collimator assembly is positioned adjacent a gamma ray acceptance surface of the detector. PA1 a. a line source having a given width and length; and PA1 b. a plurality of apertures opposite to the line source. PA1 a. providing an area detector; and PA1 b. providing a collimator at the detector that blocks gamma photons having an incident transaxial angle larger than a predetermined value. PA1 a. providing at least one area detector; PA1 b. providing at least one collimator, covering part of at least one detector; PA1 c. irradiating a patient with gamma radiation from a source positioned opposite the detector; PA1 d. collimating a flux of the gamma radiation passing through the patient from the source; PA1 e. detecting the collimated flux utilizing the portion of the area detector covered by the collimator; PA1 f. determining a two dimensional attenuation map of at least a portion of the patient from the detected flux; and PA1 g. performing a PET imaging sequence without removing the collimator. PA1 at least one detector which produces signals responsive to high and low energy events throughout a given time period; PA1 a collimator situated between a detector of the at least one detectors and the source, wherein the collimator collimates the low energy photons and is relatively transparent to the high energy photons; PA1 a dual energy detector, which receives the signals and determines therefrom whether the signal was generated by a relatively low energy photon or a relatively high energy photon; PA1 an image processing system that separately processes the high energy signals and the low energy signals to produce images based on the detected high energy and low energy photon. PA1 a first collimator which collimates low energy photons and is relatively transparent to high energy photons; and PA1 a second collimator that collimates high energy photons wherein one of the first and second collimators underlies the other of the first and second collimators overlies the other.
Another class of problems concerns the conditioning (collimation) of the radiation from a line source in transmission attenuation measurements. Reference is now made to FIGS. 4A-4C. Collimation of line source radiation is performed in one of the following ways:
Yet another problem present in coincidence measurements concerns the lack of depth discrimination due to the finite thickness of the scintillation crystal.
Reference is now made to FIGS. 5A and 5B. In coincidence measurements, a true 68 or a calculated 68' line of flight of gamma photons 70 is determined by the location of a pair of interaction points 72, of both photons in a pair 76, with detectors scintillation crystals 42. The resolution of a detector in coincidence measurements depends on:
Of these three causes, the loss of resolution in depth discrimination, (shown as X on FIG. 5B), which strongly depends on incident angle 86 and is a function of crystal's thickness 80, is most important. In order to increase depth discrimination in coincidence measurements, either the crystal thickness is reduced, or only those photons that have an angle of incidence 86, under a certain limit are counted, for example, by reducing the flux of photons with large angle of incidence.
"Septa" or "Filter" shields 56 (see FIGS. 2B, 3 and 6) have no substantial localization function per se. They only remove scattered gamma photons 60, with large axial incidence angle 86, most of which are not useful for PET and are rejected by the software. To provide this function, the septa are generally about 1 cm apart and have an acceptance angle of about 10 degrees. Prior art septa are placed parallel to the slices of the reconstructed 3D image. The limited collimation of the Septa indirectly improves resolution by reducing the effect of the lack of depth discrimination on location accuracy
It is desirable, in PET, to improve gamma detectors efficiency by reducing the number of stray photons detected relative to the number of non-stray photons detected and to improve the depth discrimination in coincidence measurements without having to reduce the scintillation crystals thickness. It is also desirable to perform attenuation and coincidence measurements in sequence without moving or replacing parts of the imaging system and, in attenuation measurements, to reduce radioactivity losses due to line source diameter while using a large diameter source to improve statistics by increasing the total radiation while keeping the source strictly collimated.